High throughput particle separation and concentration are critical for many applications in the chemical, environmental, and biomedical fields. In particular, a number of cellular and sub-cellular purification and enrichment applications are used to enable the quantitative study and diagnosis of disease. The diagnosis of infectious diseases relies on the detection of relatively small amounts of infectious organisms (e.g., viruses, bacteria, or fungi) in the blood stream or in other biological samples. It may also be desirable to isolate other components from samples, including different types of cells (e.g., cancer cells, white blood cells, etc.). Once infectious organisms or components of interest are isolated from samples, they may then be used in further applications related to identification, including the isolation of nucleic acids from those components to allow downstream processing. Thus, a need exists for methods that provide for the enrichment and purification of components of interest from biological samples, including viruses, bacteria, fungi, cancer cells, and white blood cells.
Commercial products are available for cell separation, whole blood fractionation, and subcellular fractionation using density gradients. However, these techniques have poor separation resolution and are inappropriate for recovering rare cells. Commercially available fluorescence-activated cell sorting (FACS) systems enable automated cell separation and counting, but cannot process large sample volumes and are also inappropriate for recovering rare cells.
Many microfluidic techniques have emerged for sorting, concentrating, or purifying particles and cells.1-20 Recent research into microfluidic devices has enabled applications that include microorganism recovery from environmental and biological samples21, white blood cell counting for immune deficiency diagnosis22, and circulating tumor cell (CTC) counting for cancer metastasis diagnosis and prognosis23, 24. Most of these techniques operate in the range of 1-100 μL/min, but a significantly higher throughput is needed for applications that require processing large volumes of fluid to obtain usable quantities of a target species, such as rare cell concentration. Methods designed for rare cell recovery must therefore be capable of significantly reducing the total fluid volume, from mL to 100 μL or less, to enable downstream microfluidic steps while preserving the rare targeted particles.
Recently, the need for high-throughput separation has been addressed by inertial-migration-based particle separation strategies, which are capable of achieving greater than 1 mL/min throughput.25 These strategies balance forces within the channel to locate particles of a certain size into a desired longitudinal position. In a straight channel, two inertial lift forces (FL), one due to the parabolic flow profile and the other due to the interaction of the particles with the wall, balance to focus particles to discrete equilibrium positions along the channel periphery.26,27 This was first shown by Segre and Silberberg28 at the centimeter-scale and later by others for microscale applications.29-32, 21, 33-37 
The use of curvilinear geometries such as arc,38-43 asymmetric serpentine,44-46 and spiral47-56 channels introduces a third force, the Dean force, FD, due to the formation of secondary flows, known as Dean vortices. The magnitude of these secondary flows is described by the dimensionless Dean number (De)57,58
                    De        =                                            ρ              ⁢                                                          ⁢                              U                f                            ⁢                              D                h                                      μ                    ⁢                                                    D                h                                            2                ⁢                R                                                                        (        1        )            where ρ is the density of the fluid, Uf is the average fluid velocity, μ is the fluid dynamic viscosity, R is the radius of curvature of the curvilinear channel, and Dh, is the hydraulic diameter. The hydraulic diameter, Dh, is determined from
                              D          h                =                              2            ⁢            wh                                w            +            h                                              (        2        )            where w and h are the channel width and height, respectively. The Dean vortices are counter-rotating and act to laterally displace particles across the channel by imposing a drag force. The Dean force acts in combination with the combined lift forces to alter the equilibrium positions of focused particles to a single equilibrium position near the inner wall of the channel.
FIG. 1A is a schematic diagram of a microfluidic channel cross-section illustrating the principle of inertial spiral microfluidics. The main flow (into the page) follows a curvilinear path leading to the development of secondary flows, known as Dean vortices (dashed lines). Dispersed particles experience a combination of lift forces (FL) and Dean forces (FD), which result in differential migration of the particles to unique equilibrium positions near the inner wall.
The lift forces go as FL∝ap4 and the Dean forces go as FD∝ap, where ap is particle diameter. The equilibrium position is thus particle-size-dependent, with larger particles (dominated by FL) aligning near the inner wall and smaller particles (dominated by FD) near the channel center47. An example of a two-particle separation in a microfluidic spiral is shown in FIG. 1B. Particles of one size (e.g., 15 μm) are separated from particles of a different size (e.g., 8 μm) at a flow rate of 1 mL/min.
Spiral inertial microfluidic devices have been successfully used in a wide range of applications including particle47,49,51,54 and cell49 separations, cell synchronization,50 circulating tumor cell isolation,59 and electroporation.55 These devices typically utilize a branched outlet to collect the concentrated, focused particle or cell streams. The separation efficiency for a device is defined as the number of targeted particles collected at a single outlet over the number of those particles input into the device. The concentration factor is defined as the concentration of particles, assuming 100% recovery, over the inlet partial concentration multiplied by the separation efficiency. Ultimately, for any geometry, the concentration factor is limited by the number of outlets: if there are too many, a particle stream cannot be precisely controlled to flow through a single one. Devices with branched outlets of up to eight channels have been shown49. This resulted in a concentration factor of 8× (i.e., 87.5% removal of the inlet fluid). Using only branched outlets, however, further increases in concentration factor to greater than 10× (90% fluid removal) are challenging. For example, 93% removal corresponds to a 14× concentration factor, which would require 14 outlets. The width of the outlets typically needs to be 4-5 times the particle diameter to ensure that the particle stream is collected in a single outlet. A large number of outlets, therefore, thus requires a large expansion of the width of the channel in front of the outlets. This expansion is accompanied by a corresponding increase in the width of the fluid streamlines, and thus the width of the particle stream, which becomes too wide to be collected in a single outlet channel, leading to particle loss and a decrease in separation efficiency.
Increasing particle concentration and purity in a microfluidic device by removing, or “skimming”, fluid from a main channel through microfluidic waste channels has been previously reported. Traditional skimming techniques rely on the natural formation of small particle-free regions near channel walls60-64 (as shown in FIG. 2A) or geometrical features65 (as shown at FIG. 2B). In one approach, posts66,67 or dams68 located at the outer wall were used to filter particles. Fluid and smaller particles were able to pass through the post or dam filters and were removed, while larger particles were retained. These devices were operated at low flow rates (<100 μL/min) in order to minimize the inertial effects and the secondary Dean flows. As a result, centrifugal forces dominated and pushed particles to the outside wall of the curvilinear channels. Another design took advantage of centrifugal forces to push blood cells against the outer wall and remove plasma from a waste channel on the inner wall.38 Flow rates up to 120 μL/min were achieved. In a third approach, inertial lift forces focused targeted cells away from the walls, creating a target-free region where waste channels removed fluid and non-targeted cells69,70; this design achieved >500 μL/min. These devices focused on enriching targeted cells relative to high-concentration, non-targeted species, and were not aiming to achieve high concentration factors.